Magnetic resonance imaging (MRI) is an established technology for high resolution three-dimensional imaging of the living human body, without the need for potentially hazardous x-rays. See, Joseph P. Hornak, Ph.D., The Basics of MRI, www.cis.rit.edu/htbooks/mri/(copyright 1996-2010), expressly incorporated herein by reference. This is based on magnetic resonance of the magnetic moments of hydrogen nuclei (protons) in the presence of a magnetic field. The process involves resonant absorption and emission of a radio-frequency (RF) signal in the tissue, at a frequency given by f=γB, where B is the field in teslas (T) and γ=43 MHz/T for the proton. Since the signal strength and the spatial resolution generally increase with larger fields, conventional MRI technology typically uses a field of at least B=1.5 T, corresponding to f=65 MHz. A uniform, low-noise magnetic field of this magnitude requires a superconducting magnet, which yields a large and bulky system.
MRI is sensitive to the magnetic field gradient, temperature, and microenvironment. By using a strong, uniform field, many relatively weak contributing factors are generally ignored in the analysis. However, using powerful computing devices, it is possible to analyze each voxel (minimum unit of volume in the image space) dependent on its particular magnetic field strength, gradient, and therefore draw conclusions about its microenvironment and structure, which is the typical goal of an MRI procedure.
One disadvantage of this conventional method is that it can be quite slow to achieve high resolution imaging over a substantial volume of tissue. An imaging time of many minutes can be inconvenient for the patient, as well as being inconsistent with moving tissues such as the heart and lungs. Functional MRI (fMRI) acquires a series of MRI images which are time-synchronized with a physiological status, i.e., heartbeat, to permit imaging of moving tissues, and perhaps more importantly, an analysis of their dynamic attributes.
The slow imaging time is limited by scanning magnetic field gradients in three dimensions to select a small region corresponding to the resonance condition. A single receiver coil typically measures this RF signal. One approach to accelerating the image acquisition is to use a plurality of receiver coils, with different locations, in order to obtain some spatial information and reduce the need for the full gradient scanning. A fully parallel approach which can eliminate entirely at least one dimension of gradient scanning is sometimes called “single echo acquisition”, and comprises using an array of N closed spaced (but decoupled) receiver coils, where N=64 coils has been demonstrated. (See, for example, U.S. Pat. No. 6,771,071, “Magnetic resonance imaging using a reduced number of echo acquisitions”, expressly incorporated herein by reference.) While this does increase the image acquisition speed, this approach has not yet been adopted for commercial MRI systems, due in part to the complex network of RF receivers required.
A recently developed alternative technology for MRI uses ultra-low magnetic fields, with measurement fields as low as the microtesla range, and frequencies as low as 1 kHz instead of 65 MHz. See, for example, U.S. Pat. Nos. 6,159,444; 6,885,192; 7,053,410; 7,117,102; 7,218,104; 7,187,169; and published US patent application 2009/0072828, expressly incorporated herein by reference.
This runs counter to the conventional trend of using ever higher magnetic fields to improve signal integrity and resolution. However, this approach can make use of ultra-sensitive magnetic field detectors called superconducting quantum interference devices, or SQUIDs. The sensitivity of the SQUID detector can partially compensate for the weak signal at these very low frequencies. A system of this type does not address the imaging speed issue (which may even be slower due to additional required signal averaging), but such a system would be substantially more compact and portable due to the small magnetic field, and therefore likely small magnet. Still, work to date has demonstrated only a fairly low-resolution system, which would limit its applicability.
One suggested improvement on the ultra-low field MRI system, to increase the signal-to-noise ratio and the resolution, is to use a plurality of SQUID receivers. This would also enable the same system to be used for magnetoencephalography (MEG) or magnetocardiography (MCG), where the array of SQUIDs could also localize sources of electrical currents in nerves or heart muscle. (See, for example, U.S. Pat. No. 7,573,268, “Direct imaging of neural currents using ultralow field magnetic resonance techniques”, expressly incorporated herein by reference, and V. Zotev et al, “Parallel MRI at Microtesla Fields”, J. Magnetic Resonance, vol. 192, p. 197, 2008.) However, this partially parallel approach did not address how this small array (with 7 elements) could be scaled to much larger numbers to achieve large scan-time acceleration factors.
It is well known in the art that SQUIDs are capable of ultra-sensitive measurements of magnetic fields at low frequencies of order a kilohertz or less. However, it is less well known that the SQUID itself is actually a high-frequency device, capable of field measurement up to GHz frequencies. The frequency limitation in conventional SQUID systems is actually in the external control loop that extends the dynamic range and linearity of the device as an analog sensor of magnetic field strength and/or gradient. Two complementary approaches have been developed to adapt SQUIDs to practical RF systems. In one approach, arrays of SQUIDs (sometimes called superconducting quantum interference filters or SQIFs) are coupled together to increase linearity and dynamic range, without requiring an external control loop. (See, for example, U.S. Pat. Nos. 6,690,162 and 7,369,093, expressly incorporated herein by reference.) In another approach, sometimes called a digital SQUID or a SQUID-based digitizer, the SQUID is used to generate fast single-flux-quantum (SFQ) voltage pulses, which are processed by rapid-single-flux-quantum (RSFQ) superconducting digital logic circuits to achieve a combination of linearity and dynamic range, as well as the flexibility of digital processing. See, for example, U.S. Pat. Nos. 5,420,586; 7,365,663; 7,598,897, expressly incorporated herein by reference.
The present technology provides improvements for an MRI system that may help achieve the dual goals of fine imaging, while also obtaining substantial acceleration of image acquisition. Likewise, the present technology may also assist in reducing the system bulk, and enhancing flexibility.
A useful background of MRI is found in en.wikipedia.org/wiki/Magnetic_resonance_imaging (Nov. 22, 2009), expressly incorporated herein by reference. In order to acquire a Magnetic Resonance Image, Radio frequency (RF) fields are used to periodically align a magnetic moment of a portion of hydrogen (or certain other atomic isotopes), which then relax to their unaligned state over time. Certain nuclei such as 1H (protons), 2H, 3He, 13C, 23Na or 31P, have a non-zero spin and therefore a magnetic moment. In the case of the so-called spin-½ nuclei, such as 1H, there are two spin states, sometimes referred to as “up” and “down”. When these spins are placed in a strong external magnetic field they precess around an axis along the direction of the field. Protons align in two energy “eigenstates” (the “Zeeman effect”): one low-energy and one high-energy, which are separated by a very small splitting energy.
An image may be made, on a per-voxel basis, of the magnitude and relaxation time of the magnetic alignment. The frequency at which the protons resonate depends on the strength of the magnetic field. This field-strength dependence therefore allows a frequency encoding of position. By superposing fields which predictably alter the magnetic field in known manner, different coordinates may be “addressed”, allowing full image formation, for example as a set of slices, or as a three dimensional matrix.
Different tissues can be distinguished because different chemicals, prototypically water and lipid, can be detected because the protons in different chemical compositions return to their equilibrium state at different rates. There are other aspects of magnetic resonance that can also be exploited to extract information. For example, in addition to relaxation time, local environments can also create perturbation of magnetic fields, and the presence of characteristic perturbations can be used to infer the local environment. For example, a single proton can predictably perturb the magnetic field of another proton located proximately on the same molecule.
In the static magnetic fields commonly used in MRI, the energy difference between the nuclear spin states corresponds to a radio frequency photon. Resonant absorption of energy by the protons due to an external oscillating magnetic field will occur at the Larmor frequency for the particular nucleus.
The net magnetization vector has two states. The longitudinal magnetization is due to a tiny excess of protons in the lower energy state. This gives a net polarization parallel to the external field. Application of an RF pulse can tip sideways (with i.e. a so-called 90° pulse) or even reverse (with a so-called 180° pulse) this net polarization vector.
The recovery of longitudinal magnetization is called longitudinal or T1 relaxation and occurs exponentially with a time constant T1. The loss of phase coherence in the transverse plane is called transverse or T2 relaxation. T1 is thus associated with the enthalpy of the spin system (the number of nuclei with parallel versus anti-parallel spin) while T2 is associated with its entropy (the number of nuclei in phase).
When the radio frequency pulse is turned off, the transverse vector component produces an oscillating magnetic field which induces a small current in the receiver coil. This signal is called the free induction decay (FID). In an idealized nuclear magnetic resonance experiment, the FID decays approximately exponentially with a time constant T2, but in practical MRI small differences in the static magnetic field at different spatial locations (“inhomogeneities”) cause the Larmor frequency to vary across the body creating destructive interference which shortens the FID. The time constant for the observed decay of the FID is called the T2* relaxation time, and is always shorter than T2. Also, when the radio frequency pulse is turned off, the longitudinal magnetization starts to recover exponentially with a time constant T1.
In MRI, the static magnetic field is caused to vary across the body (a field gradient), so that different spatial locations become associated with different precession frequencies. Usually these field gradients are pulsed, and a variety of RF and gradient pulse sequences may be used. Application of field gradient destroys the FID signal, but this can be recovered and measured by a refocusing gradient (to create a so-called “gradient echo”), or by a radio frequency pulse (to create a so-called “spin-echo”). The whole process can be repeated when some T1-relaxation has occurred and the thermal equilibrium of the spins has been more or less restored.
Typically in soft tissues T1 is around one second while T2 and T2* are a few tens of milliseconds, but these values vary widely between different tissues (and different external magnetic fields), permitting MRI distinguish different types of soft tissues. Contrast agents work by altering (shortening) the relaxation parameters, especially T1.
A number of schemes have been devised for combining field gradients and radio frequency excitation to create an image: 2D or 3D reconstruction from projections, much as in Computed Tomography; Building the image point-by-point or line-by-line; Gradients in the RF field rather than the static field. Although each of these schemes is occasionally used in specialist applications, the majority of MR Images today are created either by the Two-Dimensional Fourier Transform (2DFT) technique with slice selection, or by the Three-Dimensional Fourier Transform (3DFT) technique. Another name for 2DFT is spin-warp. The 3DFT technique is rather similar except that there is no slice selection and phase-encoding is performed in two separate directions. Another scheme which is sometimes used, especially in brain scanning or where images are needed very rapidly (such as in functional MRI or fMRI), is called echo-planar imaging (EPI): In this case, each RF excitation is followed by a train of gradient echoes with different spatial encoding.
Image contrast is created by differences in the strength of the NMR signal recovered from different locations within the sample. This depends upon the relative density of excited nuclei (usually water protons), on differences in relaxation times (T1, T2, and T2*) of those nuclei after the pulse sequence, and often on other parameters. Contrast in most MR images is actually a mixture of all these effects, but careful design of the imaging pulse sequence allows one contrast mechanism to be emphasized while the others are minimized. In the brain, T1-weighting causes the nerve connections of white matter to appear white, and the congregations of neurons of gray matter to appear gray, while cerebrospinal fluid (CSF) appears dark. The contrast of white matter, gray matter and cerebrospinal fluid is reversed using T2 or T2* imaging, whereas proton-density-weighted imaging provides little contrast in healthy subjects. Additionally, functional parameters such as cerebral blood flow (CBF), cerebral blood volume (CBV) or blood oxygenation can affect T1, T2 and T2* and so can be encoded with suitable pulse sequences.
In some situations it is not possible to generate enough image contrast to adequately show the anatomy or pathology of interest by adjusting the imaging parameters alone, in which case a contrast agent may be administered. This can be as simple as water, taken orally, for imaging the stomach and small bowel. However, most contrast agents used in MRI are selected for their specific magnetic properties. Most commonly, a paramagnetic contrast agent (usually a gadolinium compound) is given. Gadolinium-enhanced tissues and fluids appear extremely bright on T1-weighted images. This provides high sensitivity for detection of vascular tissues (e.g., tumors) and permits assessment of brain perfusion (e.g., in stroke).
More recently, superparamagnetic contrast agents, e.g., iron oxide nanoparticles, have become available. These agents appear very dark on T2-weighted images and may be used for liver imaging, as normal liver tissue retains the agent, but abnormal areas (e.g., scars, tumors) do not. They can also be taken orally, to improve visualization of the gastrointestinal tract, and to prevent water in the gastrointestinal tract from obscuring other organs (e.g., the pancreas). Diamagnetic agents such as barium sulfate have also been studied for potential use in the gastrointestinal tract, but are less frequently used.
In 1983 Ljunggren and Tweig independently introduced the k-space formalism, a technique that proved invaluable in unifying different MR imaging techniques. They showed that the demodulated MR signal S(t) generated by freely processing nuclear spins in the presence of a linear magnetic field gradient G equals the Fourier transform of the effective spin density. Mathematically:S(t)={tilde over (ρ)}eff({right arrow over (k)}(t))≡∫d{right arrow over (x)}ρ({right arrow over (x)})·e2πi{right arrow over (k)}(t)·{right arrow over (x)}where:{right arrow over (k)}(t)≡∫0t{right arrow over (G)}(τ)dτ
In other words, as time progresses the signal traces out a trajectory in k-space with the velocity vector of the trajectory proportional to the vector of the applied magnetic field gradient. By the term effective spin density we mean the true spin density ρ({right arrow over (x)}) corrected for the effects of T1 preparation, T2 decay, dephasing due to field inhomogeneity, flow, diffusion, etc., and any other phenomena that affect that amount of transverse magnetization available to induce signal in the RF probe. From the basic k-space formula, it follows immediately that an image I({right arrow over (x)}) may be constructed by taking the inverse Fourier transform of the sampled data:I({right arrow over (x)})=∫d{right arrow over (x)}S({right arrow over (k)}(t))·e−2πi{right arrow over (k)}(t)·{right arrow over (x)}
In a standard spin echo or gradient echo scan, where the readout (or view) gradient is constant (e.g. G), a single line of k-space is scanned per RF excitation. When the phase encoding gradient is zero, the line scanned is the kx axis. When a non-zero phase-encoding pulse is added in between the RF excitation and the commencement of the readout gradient, this line moves up or down in k-space, i.e., we scan the line ky=constant. In single-shot EPI, all of k-space is scanned in a single shot, following either a sinusoidal or zig-zag trajectory. Since alternating lines of k-space are scanned in opposite directions, this must be taken into account in the reconstruction. Multi-shot EPI and fast spin echo techniques acquire only part of k-space per excitation. In each shot, a different interleaved segment is acquired, and the shots are repeated until k-space is sufficiently well-covered. Since the data at the center of k-space represent lower spatial frequencies than the data at the edges of k-space, the TE value for the center of k-space determines the image's T2 contrast.
The importance of the center of k-space in determining image contrast can be exploited in more advanced imaging techniques. One such technique is spiral acquisition—a rotating magnetic field gradient is applied, causing the trajectory in k-space to spiral out from the center to the edge. Due to T2 and T2* decay the signal is greatest at the start of the acquisition, hence acquiring the center of k-space first improves contrast to noise ratio (CNR) when compared to conventional zig-zag acquisitions, especially in the presence of rapid movement.
Since {right arrow over (x)} and {right arrow over (k)} are conjugate variables (with respect to the Fourier transform) we can use the Nyquist theorem to show that the step in k-space determines the field of view of the image (maximum frequency that is correctly sampled) and the maximum value of k sampled determines the resolution for each axis, i.e.,
  FOV  ∝            1              Δ        ⁢                                  ⁢        k              ⁢                  ⁢    Resolution    ∝                          k        max                    .  
In acquiring a typical MRI image, Radio frequencies are transmitted at the Larmor frequency of the nuclide to be imaged. For example, for 1H in a magnetic field of 1 T, a frequency of 42.5781 MHz would be employed. During the first part of the pulse sequence, a shaped pulse, e.g., using sinc modulation, causes a 90° nutation of longitudinal nuclear magnetization within a slab, or slice, creating transverse magnetization. During the second part of the pulse sequence, a phase shift is imparted upon the slice-selected nuclear magnetization, varying with its location in the Y direction. During the third part of the pulse sequence, another slice selection (of the same slice) uses another shaped pulse to cause a 180° rotation of transverse nuclear magnetization within the slice. This transverse magnetization refocuses to form a spin echo at a time TE. During the spin echo, a frequency-encoding (FE) or readout gradient is applied, making the resonant frequency of the nuclear magnetization vary with its location in the X direction. The signal is sampled nFE times by the ADC during this period. Typically nFE of between 128 and 512 samples are taken. The longitudinal magnetization is then allowed to recover somewhat and after a time TR the whole sequence is repeated nPE times, but with the phase-encoding gradient incremented. Typically nPE of between 128 and 512 repetitions are made. Negative-going lobes in GX and GZ are imposed to ensure that, at time TE (the spin echo maximum), phase only encodes spatial location in the Y direction. Typically TE is between 5 ms and 100 ms, while TR is between 100 ms and 2000 ms.
After the two-dimensional matrix (typical dimension between 128×128 and 512×512) has been acquired, producing the so-called K-space data, a two-dimensional Fourier transform is performed to provide the familiar MR image. Either the magnitude or phase of the Fourier transform can be taken, the former being far more common.
The magnet is a large and expensive component of an MRI scanner. The strength of the magnet is measured in tesla (T). Clinical magnets generally have a field strength in the range 0.1-3.0 T, with research systems available up to 9.4 T for human use and 21 T for animal systems. Just as important as the strength of the main magnet is its precision. The straightness of the magnetic lines within the center (or, as it is technically known, the iso-center) of the magnet needs to be near-perfect. This is known as homogeneity. Fluctuations (inhomogeneities in the field strength) within the scan region should be less than three parts per million (3 ppm). Three types of magnets have been used:
Permanent magnet: Conventional magnets made from ferromagnetic materials (e.g., iron alloys or compounds containing rare earth elements such as neodymium) can be used to provide the static magnetic field. A permanent magnet that is powerful enough to be used in a traditional type of MRI will be extremely large and bulky; they can weigh over 100 tons. Permanent magnet MRIs are very inexpensive to maintain; this cannot be said of the other types of MRI magnets, but there are significant drawbacks to using permanent magnets. They are only capable of achieving weak field strengths compared to other MRI magnets (usually less than 0.4 T) and they are of limited precision and stability. Permanent magnets also present special safety issues; since their magnetic fields cannot be “turned off,” ferromagnetic objects are virtually impossible to remove from them once they come into direct contact. Permanent magnets also require special care when they are being brought to their site of installation.
Resistive electromagnet: A solenoid wound from copper wire is an alternative to a permanent magnet. An advantage is low initial cost, but field strength and stability are limited. The electromagnet requires considerable electrical energy during operation which can make it expensive to operate. This design is considered obsolete for typical application.
Superconducting electromagnet: When a niobium-titanium or niobium-tin alloy is cooled by liquid helium to 4 K (−269° C., −452° F.) it becomes a superconductor, losing resistance to flow of electrical current. An electromagnet constructed with superconductors can have extremely high field strengths, with very high stability. The construction of such magnets is costly, and the cryogenic helium imposes operating costs and requires careful handling. However, despite their cost, helium cooled superconducting magnets are the most common type found in MRI scanners today.
Magnetic field strength is an important factor in determining image quality. Higher magnetic fields increase signal-to-noise ratio, permitting higher resolution or faster scanning. However, higher field strengths require more costly magnets with higher maintenance costs, and have increased safety concerns. A field strength of 1.0-1.5 T is considered a good compromise between cost and performance for general medical use. However, for certain specialist uses (e.g., brain imaging) higher field strengths are desirable, with some hospitals now using 3.0 T scanners.
Gradient coils are used to spatially encode the positions of protons by varying the magnetic field linearly across the imaging volume. The Larmor frequency will then vary as a function of position in the x, y and z-axes. Gradient coils are usually resistive electromagnets powered which permit rapid and precise adjustments to their field strength and direction. Typical gradient systems are capable of producing gradients from 20 mT/m to 100 mT/m (i.e., in a 1.5 T magnet, when a maximal Z-axis gradient is applied, the field strength may be 1.45 T at one end of a 1 m long bore and 1.55 T at the other). It is the magnetic gradients that determine the plane of imaging—because the orthogonal gradients can be combined freely, any plane can be selected for imaging. Scan speed is typically dependent on performance of the gradient system. Stronger gradients allow for faster imaging, or for higher resolution; similarly, gradients systems capable of faster switching can also permit faster scanning. However, gradient performance may also be limited by safety concerns over possible nerve stimulation.
The radio frequency (RF) transmission system consists of an RF synthesizer, power amplifier and transmitting coil. The receiver in a traditional system consists of a receiving coil, pre-amplifier and signal processing system. It is possible to scan using an integrated coil for RF transmission and MR signal reception, but if a small region is being imaged, then better image quality (i.e. higher signal-to-noise ratio) is obtained by using a close-fitting smaller coil.
A recent development in MRI technology has been the development of sophisticated multi-element phased array coils which are capable of acquiring multiple channels of data in parallel. See, Roemer P B, Edelstein W A, Hayes C E, Souza S P, Mueller O M (1990). “The NMR phased array”. Magnetic Resonance in Medicine 16 (2): 192-225. This ‘parallel imaging’ technique uses acquisition schemes that allow for accelerated imaging, by replacing some of the spatial coding originating from the magnetic gradients with the spatial sensitivity of the different coil elements. However, the increased acceleration also reduces the signal-to-noise ratio and can create residual artifacts in the image reconstruction. Two frequently used parallel acquisition and reconstruction schemes are known as SENSE and GRAPPA. See, Pruessmann K P, Weiger M, Scheidegger M B, Boesiger P (1999). “SENSE: Sensitivity encoding for fast MRI”. Magnetic Resonance in Medicine 42 (5): 952-962. doi:10.1002/(SICI)1522-2594(199911)42:5<952::AID-MRM16>3.0.CO;2-S. PMID 10542355, Griswold M A, Jakob P M, Heidemann R M, Nittka M, Jellus V, Wang J, Kiefer B, Haase A (2002). “Generalized autocalibrating partially parallel acquisitions (GRAPPA)”. Magnetic Resonance in Medicine 47 (6): 1202-1210. doi:10.1002/mrm.10171. PMID 12111967; Blaimer M, Breuer F, Mueller M, Heidemann R M, Griswold M A, Jakob P M (2004). “SMASH, SENSE, PILS, GRAPPA: How to Choose the Optimal Method”. Topics in Magnetic Resonance Imaging 15 (4): 223-236. cfmriweb.ucsd.edu/ttliu/be280a_05/blaimer05.pdf.
Hydrogen is the most frequently imaged nucleus in MRI because it is present in biological tissues in great abundance. However, any nucleus which has a net nuclear spin could potentially be imaged with MRI. Such nuclei include helium-3, carbon-13, fluorine-19, oxygen-17, sodium-23, phosphorus-31 and xenon-129. 23Na and 31P are naturally abundant in the body, so can be imaged directly. Gaseous isotopes such as 3He or 129Xe must be hyperpolarized and then inhaled as their nuclear density is too low to yield a useful signal under normal conditions. 17O, 13C and 19F can be administered in sufficient quantities in liquid form (e.g. 17O-water, 13C-glucose solutions or perfluorocarbons) that hyperpolarization is not a necessity.
Multinuclear imaging is primarily a research technique at present. However, potential applications include functional imaging and imaging of organs poorly seen on 1H MRI (e.g. lungs and bones) or as alternative contrast agents. Inhaled hyperpolarized 3He can be used to image the distribution of air spaces within the lungs. Injectable solutions containing 13C or stabilized bubbles of hyperpolarized 129Xe have been studied as contrast agents for angiography and perfusion imaging. 31P can potentially provide information on bone density and structure, as well as functional imaging of the brain.
Portable magnetic resonance instruments are available for use in education and field research. Using the principles of Earth's field NMR, they have no powerful polarizing magnet, so that such instruments can be small and inexpensive. Some can be used for both EFNMR spectroscopy and MRI imaging, e.g., the Terranova-MRI Earth's Field MRI teaching system (www.magritek.com/terranova.html). The low strength of the Earth's field results in poor signal to noise ratios, requiring long scan times to capture spectroscopic data or build up MRI images.
A magnetic resonance system that can flexibly combine a plurality of signals and antennas, permitting faster generation of high-resolution medical images in a safe and economical system, is desirable.